1. Field of the Invention
The present invention relates to, inter alia, medical imaging systems, and, in particular, to techniques for correcting misalignment effects in reconstructed images, such as, e.g., for misaligned gamma cameras of nuclear medicine imaging systems and/or the like.
2. Background Discussion
1. General Background
A variety of medical imaging systems are known. Some illustrative imaging systems include nuclear medical imaging systems (e.g., gamma cameras), computed tomography (CT or CAT) systems, magnetic resonance imaging (MRI) systems, positron-emission tomography (PET) systems, ultrasound systems and/or the like.
With respect to nuclear medical imaging systems, nuclear medicine is a unique medical specialty wherein radiation (e.g., gamma radiation) is used to acquire images that show the function and/or anatomy of organs, bones and/or tissues of the body. Typically, radioactive compounds, called radiopharmaceuticals or tracers, are introduced into the body, either by injection or ingestion, and are attracted to specific organs, bones or tissues of interest. These radiopharmaceuticals produce gamma photon emissions that emanate from the body and are absorbed by a scintillation crystal, which produces flashes of light or “events.” These events can be detected by an array of photo-detectors, such as photomultiplier tubes, and their spatial locations or positions can be calculated and stored. In this manner, an image of an organ, tissue or the like under study can be created from the detection of the distribution of the radioisotopes in the body. Typically, one or more detectors are used to detect the emitted gamma photons, and the information collected from the detector(s) is processed to calculate the position of origin of the emitted photon from the source (i.e., the body organ or tissue under study). The accumulation of a large number of emitted gamma positions allows an image of the organ or tissue under study to be displayed.
FIG. 1 depicts components of a typical nuclear medical imaging system 100 (i.e., having a gamma or scintillation camera) which includes a gantry 102 supporting one or more detectors 108 enclosed within a metal housing and movably supported proximate a patient 106 located on a patient support (e.g., pallet or table) 104. Typically, the positions of the detectors 108 can be changed to a variety of orientations to obtain images of a patient's body from various angles and locations along the patient's body. In many instances, a data acquisition console 200 (e.g., with a user interface and/or display) is located proximate a patient during use for a technologist 107 to manipulate during data acquisition. In addition to the data acquisition console 200, images are often “reconstructed” or developed from the acquired image data (“projection data”) via a processing computer system which is operated at another image processing computer console including, e.g., an operator interface and a display, which may often be located in another room, to develop images. By way of example, the image acquisition data may, in some instances, be transmitted to the processing computer system after acquisition using the acquisition console.
Nuclear medicine imaging typically involves the assessment of a radionuclide distribution within a patient after the in vivo administration of radiopharmaceuticals. Imaging systems that assess radionuclide distribution include radiation detectors and acquisition electronics. Typically, the imaging systems detect x-ray or gamma ray photons derived from the administered radionuclides. Single photon emission imaging and coincidence imaging are two forms of nuclear medicine imaging that are currently in common use. In single photon emission imaging, the radionuclide itself directly emits the radiation to be assessed. For example, in Single Photon Emission Computed Tomography (SPECT), γ-emitting radionuclides such as 99mTc, 123I, 67Ga and 111In may be part of the administered radiopharmaceutical.
Detectors used in such single photon emission imaging often use collimators placed between the patient and the gamma ray camera of the detector. In general, the collimators help to eliminate substantially all photons but those photons traveling in a desired direction from impinging on the detector surface. This is desirable to prevent scattered or background radiation photons from spuriously contributing to the image, and thereby causing inaccuracies. For example, a parallel hole collimator helps to eliminate from detection photons traveling in all directions except those almost perpendicular to the surface of the detector. The energy of emitted photons as well as their location of origin may then be accumulated until a satisfactory amount of projection data is acquired to allow the reconstruction of a clinically significant image.
Coincidence imaging helps to eliminate the need for such a collimator by relying on the detection of two oppositely traveling gamma photons, emitted as a result of the annihilation of a positron, at oppositely located detectors at nearly the same time. An example of coincidence imaging in current clinical use is Positron Emission Tomography (PET).
Typically, radiation detectors used in nuclear medicine imaging need to absorb x- or gamma-ray photons in an energy range typically between 1 keV and several MeV. These imaging photons are the photons either directly emitted or resulting from radionuclides within a patient. In order to stop spurious photons of similar energies with a collimator in SPECT imaging, a material with a high density and a high atomic number (Z) is necessary. Lead is the most common material used for collimators, but other materials such as, e.g., tungsten may also be used.
Typically, in radiology, detectors used clinically only integrate the energy deposited by a beam. However, a new generation of detectors for digital radiography and computed tomography (CT) can obtain extra information by counting individual photons and measuring their energy.
With respect to scintillators, a variety of scintillators are known. For example, scintillators include, e.g., continuous single slab, pixilated and/or columnar grow crystals. As for radionuclide imagers with pixilated radiation detector elements, typically cadmium zinc telluride (“CZT”) crystals have recently been developed. In these pixilated radionuclide imagers, the intrinsic spatial resolution is defined by the size of the individual pixilated detector elements, rather than the separation between collimator holes. See, e.g., U.S. Pat. No. 6,838,672, assigned to the present assignee, the entire disclosure of which is incorporated herein by reference. With respect to the use of CZT as a solid state (i.e., semiconductor) detector material, as a single photon detector, CZT is typically superior to NaI in several performance parameters. Among other things, the count rate capability for CZT detectors is virtually unlimited as compared to a typical scintillator crystal, because each pixel (or picture element) of the CZT material can act as an independent detector. Thus, unlike a typical scintillator crystal, in which two events occurring very close in time and spatial location will produce overlapping light output, two gamma photons arriving at exactly the same time in adjacent pixels of a CZT detector could be independently detected and measured accurately with respect to energy, given an optimum electronic circuitry design.
2. Gantry Misalignments
Current tomographic reconstructions rely on center-of-rotation (COR) and multi-head-registration (MHR) calibration schemes to correct for errors in the alignments of gantry/detector systems. These schemes are capable of removing the effects of certain misalignments (e.g., pure translational offsets and the average effects of angular misalignments). These “correctable” components are corrected by shifting the projection data before or during reconstruction and by correcting the angle at which the data are back projected.
Other components of resolution loss cannot be removed by simple manipulation of the projection data because the misalignment causes blurring in the axial dimension. That is, the image of a point source oscillates axially in the projection images and the oscillation depends on the location of the source within the object.
For gantries with cantilevered heads, this is a particular problem because of head droop. Head droop causes significant non-correctable axial blurring, which not only distorts sagital and coronal images, but reduces contrast in transaxial images (because the counts are misplaced into other slices). In addition, reducing head droop via more rigid mechanics is expensive.
While a variety of background technologies exist, there is a continued need in the art for improved systems and methods for, among other things, accommodating for misalignments in detectors.